Abstract

This article reports on a retinal stimulation system for long-term use in animal electrical stimulation experiments. The presented system consisted of an implantable stimulator which provided continuous electrical stimulation, and an external component which provided preset stimulation patterns and power to the implanted stimulator via a paired radio frequency (RF) coil. A rechargeable internal battery and a parameter memory component were introduced to the implanted retinal stimulator. As a result, the external component was not necessary during the stimulation mode. The inductive coil pair was used to pass the parameter data and to recharge the battery. A switch circuit was used to separate the stimulation mode from the battery recharging mode. The implantable stimulator was implemented with IC chips and the electronics, except for the stimulation electrodes, were hermetically packaged in a biocompatible metal case. A polyimide-based gold electrode array was used. Surgical implantation into rabbits was performed to verify the functionality and safety of this newly designed system. The electrodes were implanted in the suprachoroidal space. Evoked cortical potentials were recorded during electrical stimulation of the retina. Long-term follow-up using OCT showed no chorioretinal abnormality after implantation of the electrodes.

1. Introduction

Retinal prostheses are under investigation by several groups [1, 2], and some
preclinical and clinical trials have been reported [3–6]. Preclinical experiments
are intended to estimate the stimulation parameters and to evaluate the
efficacy and the safety of the devices and the clinical trials are aimed at
demonstrating the feasibility of the prosthesis. Although the electrical
stimulation of the retina in preliminary clinical studies showed encouraging
results such as the patient's perception of a small spot of light or basic
shapes according to the stimulation pattern [4], there is much to be
investigated and revised regarding retinal prostheses.

Before the implantation of a retinal prosthesis into the eye of a patient, the
stimulation conditions, the long-term stability, and the durability of the
retinal prostheses should be verified in vivo. Animal experiments are
essential in developing retinal prosthesis because the characteristics and
safety of prosthesis cannot be verified in human eyes for ethical reasons. Due to anatomical
differences, devices for animal use cannot be used in humans, but design
features can be tested in extended animal experiments, verifying the design
features for a human device.

Previous reports showed long-term biocompatibility of various electrodes without electrical
stimulation of the retina [7–10]. Thresholds of electrical stimulation were
presented from acute implantation experiments [11–13] and also from chronic
human studies [14, 15]. In previous studies, electrical stimulation systems
needed a wired or a wireless external component to provide stimulation
parameters and power to the internal electrodes. When those systems were used
for long-term electrical stimulation in animals [16, 17], there were many
problems that needed to be solved. Systems with a percutaneous connection to
the external portion restricted the animals' movement and posed an infection
risk. In systems having transcutaneous connections to an external controller
worn by the animals, the external part usually separated from the animals or
was damaged.

In this article, an implantable retinal prosthesis system is proposed for a chronic electrical stimulation test
in an animal model. For this purpose, a small rechargeable battery and a
parameter memory were introduced into the implanted stimulator so the external
power supply and control part could be removed during a chronic stimulation
experiment. The animal is then free of any external components. The external
unit is then needed temporarily for two purposes only: passing the parameter
and charging the battery. To our knowledge, this is the only retina stimulation
system designed for use in animals with such a feature.

Animal experiments were done to show the
feasibility of the implantation of this newly suggested stimulation system. To
check whether the implanted electrode could induce appropriate cortical
response upon electrical stimulation, we measured the electrically evoked
cortical potentials (EECPs) from rabbits in which electrodes were implanted.
The long-term biocompatibility of electrodes was also evaluated in vivo with OCT.

2. Methods

2.1. Retinal Prosthesis System Design

The implantable retinal prosthesis system for a chronic animal experiment consisted
of an internal unit for retinal stimulation and an external unit for
stimulation control and battery charging (see Figure 1). A paired RF coil links
these two units for the transmission of data and power.

Figure 1: System block diagram.

The external unit had a stimulation waveform parameter selector used to control the
channel selection, amplitude, duration, and rate of stimulation. This parameter
selector generated a parameter data frame and was implemented using a
commercially available digital signal processing chip (TMS320VC5509, Texas
Instruments, Dallas, Tex, USA). The control codes were implemented in-house
using the C programming language, and the parameter data frame consisted of 22 bits
as shown in Figure 2. The same stimulation waveform was simultaneously
delivered to all selected channels. To transmit this parameter data into the
internal stimulator, the pulse width modulation (PWM) encoding method was used
(see Figure 2). Logic “1” and “0” were encoded to have a duty cycle of 75% and
25%, respectively, and the “end-of-frame (EOF)” bit had a 50% duty cycle. Such
an encoding method enables easier synchronization and decoding because each bit
had a uniform rising edge at its beginning [18]. The transmission data rate was
125 kbps. For transmission of PWM encoded data, a class-E tuned power amplifier
(data/power transmitter) was used with amplitude shifted keying (ASK)
modulation. The carrier frequency was 2.5 MHz.

Figure 2: Data formats and PWM modulation.

The transmitted data were received by the internal coil and then the envelope was
extracted through a half-wave rectifier and a lowpass filter. Using this
envelope signal, a data decoder in the data/power receiver chip recovered the
parameter data and saved it to the parameter memory chip. Using the same envelope
signal, a voltage regulator generated power to be consumed by the data/power
receiver chip (see Figure 1).

The internal unit, that is, the retinal stimulator, was implemented with a
rechargeable battery and integrated circuit (IC) chips including the data/power
receiver chip, current stimulation chip, parameter memory chips, and battery
charging chip. The data/power receiver chip had data decoding and voltage
regulation function blocks. Both the
data/power receiver chip and the current stimulation chip are custom IC's
designed by our laboratory (0.8 m complementary metal-oxide semiconductor (CMOS)
technology, Austria Microsystems, Unterpremstaetten,
Austria). The parameter memory and battery charging chips were off-the-shelf
commercial products (see Figure 1). Except for the data/power receiver chip,
the other chips in the stimulator were powered by a rechargeable battery.
Therefore, once the parameters were passed to the parameter memory, the
external unit can be removed from the animal during the electrical stimulation
test. The retinal stimulator had two modes of function: a stimulation mode and
a battery recharging mode.

In stimulation mode, the saved parameter data in the parameter memory component were provided to the current
stimulation chip. The parameter memory component was composed of three 8-bit
shift registers (SN54AHC595, Texas Instruments, Dallas, Tex, USA) for 22-bit
data storing. The parameter data did not change unless a new parameter was
transmitted from the external component. The current stimulation chip consisted
of seven current sources and a timing logic circuitry. The current generator
circuitry had current bias circuitry (8 μA) and an 8-bit
binary current-weighted DAC (digital-to-analog converter). The timing logic
circuitry had a 2.5 MHz oscillator and switch control logic circuitry
for controlling the current stimulation waveform.

In the battery recharging mode,
2.5 MHz sinusoidal waves were transmitted with no
data. A rechargeable coin-type lithium ion battery (PD2320, Korea Power cell Inc.,
Daejeon, Korea) was used as the power source for the internal implant. A
charging chip (LTC4054L, Linear Technology Corporation, Milpitas, Calif, USA) was
used to control the battery recharging.

In this work, only one inductive coupling was used for data transmission and battery charging. Simultaneous
transmission of the stimulation parameter and charging power is difficult
because the battery charger circuit affects the precisely designed load value
of the data/power receiving circuit and can induce the failure in data
reception. To separate the stimulation mode and battery charging mode, a switch
circuit was positioned between the voltage regulator in the data/power receiver
chip and the battery charge chip to control the recharging of the battery. The
introduced switch circuit consisted of two p-MOS transistors, one capacitor and
one resistor (see Figure 3). The resistor and capacitor comprised a parallel
connection with an RC time constant of 100 millisecons, which was very long
compared to the 8-microsecond period of the clock of the data/power receiver
chip. Therefore, the voltage of the “a” node was higher than the threshold of
the Q2 switch when a data signal (PWM) was applied causing the Q2 to turn off.
The data decoding could therefore be successfully carried out with no load
effect. However, when only a sinusoidal wave without data was applied to the
retinal stimulator through the inductive link, the level of CLK was logic high
(see Figure 3). In this case, Q1 would be turned off and the voltage of node
“a” would be logic low, so Q2 would be turned on. Therefore, the battery could be recharged.

Figure 3: Operation of switch circuitry in stimulation and battery charging modes. When no data were loaded on 2.5 MHz carrier, CLK was logic high, therefore switch Q1 would be turned off, and the
voltage of “a” node would be logic low, so switch Q2 would be turned on (battery recharging mode). If any data were loaded
on 2.5 MHz carrier, the CLK would recover and the voltage of node “a” never went
below the threshold of switch Q2, so Q2 was turned off, this would be disabling
the charger chip period (stimulation mode).

To protect the ICs from body fluids and mechanical forces, the electronics of the
stimulator were hermetically housed in a metal package which consists of
biocompatible titanium housing, platinum feedthroughs, and a ceramic plate. The
feedthroughs connected the electrode array and receiver coil to the retinal
stimulator. A ceramic sintering process was used to fix the feedthroughs in the
ceramic plate that provided electrical isolation. Brazing and laser welding
techniques were employed to achieve hermetic sealing of the titanium housing [18].

A polyimide-based seven-channel, strip-shaped (750 300 m) gold electrode array
was used as stimulating electrode. The stimulation sites were constructed in a
4 mm 4 mm area with a seven segmented configuration [17].
A large circle electrode in 1500 m diameter, also polyimide-based, was used as
the reference electrode (see Figure 4). The thickness of polyimide/gold
electrode array was 58 m. One side of this array was lengthened and connected
to the stimulator via feedthroughs.

2.2. Surgical Procedures for Implantation of the Prosthesis System into Rabbits

New Zealand White rabbits weighing 2.02.5 kg were used to evaluate the proposed system. All
procedures of animal experimentation were approved by the Institutional Review
Board of Seoul National University Hospital Clinical Research Institute and
followed the Association for Research in Vision and Ophthalmology (ARVO)
Statement on Use of Animals in Ophthalmic and Vision Research. Implantation of
the entire system was performed under general anesthesia achieved by repetitive
intramuscular injection of 25 mg ketamine and 6 mg xylazine per kilogram of body weight.

The skin was prepared and a longitudinal incision was made at the right
auscultation triangle of the back. Meticulous dissection was done between the
subcutaneous and muscular layers. A subcutaneous tunnel was made by an elevator
from the medial angle of the scapula to the inferior conjunctival fornix. The
forniceal opening was created with a blade and the tip of the elevator was
extruded from the forniceal opening thereby completing the subcutaneous tunnel.
The internal stimulator was inserted into the subcutaneous tunnel from the
opening of the back. The connection part and the polyimide electrode array were
protected by a soft polyethylene tube and passed through the tunnel. After
being introduced through the conjuctival forniceal opening, the connection wire
was turned around the eyeball under the extraocular muscles and temporarily
anchored onto the sclera with two 6-0 Prolene sutures, which is a process
similar to Humayun's method [4].

A 5 mm sized Scleral tunnel incision parallel to the limbus was made with a blade and
scissors about 5 mm away from the limbus in the upper lateral quadrant of the
eyeball. To prevent vitreous prolapse and to lower the intraocular pressure,
0.1–0.2 mL of aqueous humor was drained from the
anterior chamber before the scleral opening was created. The polyimide
electrode array was inserted through the scleral opening and slided into the
suprachoroidal space to reach the visual streak. Figure 5 shows the position of
the electrode array in the rabbit eye ball. The scleral opening was closed with
8-0 Vicryl sutures and the connection wire was permanently anchored onto the
sclera with sutures. The reference electrode was left outside the eyeball to
contact the sclera without fixation. The conjunctival incisions were repaired
with 8-0 Vicryl sutures and the skin incision was repaired with 6-0 Catgut
sutures. After the operation, the entire system was inside the body and was not
exposed.

Figure 5: A diagram showing the positioning of an electrode array in the rabbit eye ball.
The electrode array is colored in yellow. (Figure 9(a) shows a fundus photography of rabbit retina
with a suprachoroidal electrode.)

Figure 6: The voltage drop between a strip-shaped electrode (in the suprachoroidal
space under visual streak) and the reference electrode (on the surface of sclera) with 104 uA
current intensity (upon implantation.)

2.3. Recording of Visually/Electrically Evoked Cortical Potentials

Stainless needle electrodes were used as recording electrodes. An active recording
electrode was placed into the primary visual cortex 6 mm lateral and 6 mm
anterior to lambda, which is the same location as in Okuno's method [19]. A
reference recording electrode was placed into the cortex 20 mm anterior to lambda.
The animal was grounded by an electrode on the ipsilateral ear. The visually evoked
cortical potentials (VECPs) and electrically evoked cortical potentials (EECPs)
elicited by stimulating one eye were recorded from the active recording
electrode placed on the contralateral side.

An integrated hardware/software platform (TDT System 3, Tucker-Davis Technologies,
Alachua, Fla, USA) was used for amplifying, acquisition, and storage of VECP
and EECP signals. This platform could flexibly integrate devices for the
intended purpose. In our recording system, there were a four-channel
low-impedance headstage (RA4L1), a 16-channel preamplifier (RA16PA Medusa
PreAmps), a DSP device (RA16BA Medusa Base Station), and a PC interface module
(FI5/PI5-to-zBus). Signals were digitized at 25 kHz on the preamplifier and
sent over a fiber optic link to a DSP device where they were filtered (0.5–300 Hz)
and processed in real time. Ordinarily, 30 consecutive responses were summed
and averaged on one VECP/EECP record.

For recording the EECP, a cathodic-first biphasic constant current stimulus waveform was
simultaneously applied to all selected active channels. Both pulse duration and
interpulse delay were 1 millisecond and amplitudes were varied as noted. The
repetition frequency was 4 Hz.

During the electrical stimulation, real-time signals also were recorded by a cornea
contact electrode through another recording channel in the above recording
system. From these recorded signals, the stimulation artifact signals were extracted
and used as the trigger signal for EECP recording. In our measurements, the
delay of the extracted trigger signal compared to the stimulation signal was
less than 500 microseconds and had a constant value.

3. Results

3.1. Retinal Prosthesis System

A current mode, charge-balanced, cathodic-first biphasic stimulation waveform was
generated in the stimulation mode and provided to a stimulation electrode
array. The current stimulation chip could simultaneously deliver a stable
current from 8 μA to 2 μA to all channels. The pulse width and the
interphase delay could be changed up to 3 milliseconds. The interphase delay
was designed to have the same time duration as the pulse width.

The fabricated polyimide electrode array was checked for electrode impedance in vitrousing a commercial potentiostat (Zahner Elektrik IM6e, Germany).
Impedance for the strip-shaped electrode site (750 μm 300 μm) was typically
1.3 and for the reference electrode was typically
300 Ω in phosphate-buffered
solution (pH 7.4) as measured at 1 kHz.

In vivo electrode impedance of the stimulation
electrode site was checked for two weeks postsurgery. The potentiostat was
connected to the sites through a percutaneous connector. Impedance between a
strip-shaped site and the reference electrode was typically near 10 at 1 KHz and there was no significant change
during the two-weeks monitoring period. The result is detailed in Table 1.

Table 1: Measured impedance at 1 KHz in vivo
for two weeks after implantation. The strip-shaped seven-segment electrode array was
placed into the suprachoroidal space and reference electrode was placed on the surface of sclera. A
strip-shaped electrode site was connected to the working electrode input, and the
stimulating reference electrode served as the counter and reference electrode. 0 week means the day
of implantation.

The surface of the gold stimulation site was checked with an atomic force microscopy
(AFM) image (AFM, PSIA, XE-150) as shown in Figure 7. The surface average
roughness of gold electrode array was 68 nm.

Figure 7: An AFM image of gold/polyimide electrode surface. The average surface
roughness was 68 nm.

The stimulation outputs were connected to 1.3 resistive loads, modeling the electrodes. The
stimulator consumed around 2 mW when delivering 520-μA biphasic current
pulse with 1-millisecond pulse width at a stimulation rate of 4 Hz.

The capacity of the battery was 75 mAh (4.2 V) and the
battery could supply the power to the internal circuitries for over 30 hours
under the 520-μA amplitude, 1-millisecond
stimulation duration. The battery was fully recharged within three hours with
25 mA charging current through the RF inductive link
when in the battery recharging mode.

3.2. Postoperative State of Rabbits

Implantation of the stimulation system into the rabbit
was successfully achieved. The internal portions were harbored safely and postoperatively
remained in situ, as verified with fundus photography.
There were no limitations in eye movements or shoulder movements on the
implanted side and the entire stimulation system could be safely protected
under the skin during the follow-up period. There was no need to anesthetize
the rabbit while changing the stimulation parameters by attaching the external
RF coil onto the skin overlaying the internal RF coil.

3.3. VECP & EECP

Figure 8 shows a typical VECP and EECP recording for a
single rabbit with varying stimulus current. The overall shapes and latencies
of VECP waves were similar to previous reports [19]. The EECP were well
recorded upon electrical stimulation from the implanted electrodes. The EECP
disappeared after the optic nerves were severed.

Threshold current necessary to elicit the EECP's were also recorded from another animal
for six weeks postsurgery. Seven channels were stimulated simultaneously using
a biphasic pulse repeated at 4 Hz with a 1-millisecond duration/phase and a 1-millisecond
interphase delay. The threshold ranged from 80 μA to 128 μA
during the testing period but did not show any significant change. The detailed
measured threshold currents and calculated charge density values are shown in
Table 3. The threshold was the highest (128 μA)
on the day of implantation but then fell and remained at lower values for the
duration of monitoring. The threshold charge density was similar to the value
reported in [20].

Table 3: Threshold current to elicit the EECP. Seven channels were used
simultaneously and the duration of the pulse was a biphasic current pulse, 1 ms/phase with a 1-millisecond interphase
delay. The repetition rate was 4 Hz. 0 week means the day of implantation.

3.4. Long-Term Follow-up Results

The fundus photography and OCT findings are shown in
Figure 9. The Stratus OCT 3000 showed fine resolution of cross-sectional
retinal images of rabbits. We can distinguish each layer of retina and choroid
by OCT. All the implanted electrodes were detected in the suprachoroidal space,
which was in the same location of the retina as immediately after surgery. The
eyes with polyimide-based gold electrodes showed no chorioretinal reaction
around the electrodes (see Figure 9(a)). The implanted electrodes did not
cause any chorioretinal inflammation and structural deformity during 16 months
of the follow-up period.

Figure 9: (a) A digital fundus photography of rabbit retina implanted with a
suprachoroidal electrode with two OCT (Stratus OCT 3000) scan paths (solid and
dashed lines). A polyimide-based gold electrode had been implanted. Under gross
inspection, the electrode is in its original position near the visual streak
and no chorioretinal abnormalities can be found. (b) The OCT scan of the retina
remote from the electrode (the solid line in the fundus photography). The
normal structures of the rabbit retina and choroid is demonstrated. (c) The OCT
scan passing the electrode vertically (the dashed line in the fundus
photography). The electrode is observed in the suprachoroidal space and in tight
contact with the choroid. The metal plates in the electrode are highly
reflective and leave posterior shadowing. The overlying retina showed normal
structures of layers.

4. Discussion

In retinal prosthesis research, long-term electrical stimulation experiments in
animal are needed to verify the long-term stability and durability of the
system before clinical use. Retinal prosthesis systems for electrical
stimulation usually consist of external and internal units [21, 22]. The
external unit is necessary to change the stimulation parameters and to provide
power for the implanted stimulator, but it is also burdensome especially in
long-term animal electrical stimulation experiments.

The stimulation system presented in this article was intended to provide a useful
tool for long-term animal experiments on retinal prostheses. To remove the
external connection or the external unit from the animal during electrical
stimulation, we used a small rechargeable battery in the implantable
stimulator. This battery could be simply recharged using an RF inductive link
while the system was idle. This system makes it possible to conduct chronic
electrical stimulation tests in such a way that the animal can move and act
freely without any external unit restriction during the stimulation test.
Therefore, there is no need to anesthetize the animal frequently and the
stimulation system is also protected from the animal's claws and teeth.

The implantable retinal stimulator was built using IC chips and discrete elements for this
proof-of-concept. An IC chip can be developed to reduce power consumption and
further miniaturize the implanted component of the device.

During the implantation of the polyimide electrode array into the eye and the inside
of orbit, we adopted Humayun's method [4], making one turn of the electrode
along the equator of the eye under the extraocular muscles to provide
stability. The elongated polyimide electrode is flexible, thus simple folding
of the connection part can provide much degrees of freedom, and this enables
eye movement without limitation of motion.

There are pros and cons regarding the ocular location of implantation of the
electrodes. To our thought, the suprachoroidal implantation of electrodes has
many advantages over the conventional subretinal implantation.

First, as the electrode does not contact retinal cells directly, there is no risk of
mechanical retinal damage from direct contact with electrode which could be
initiated during surgical procedure. A past report comparing subretinal and
suprachoroidal retinal implantation in rabbits also showed histologically
proven retinal damage in the subretinal implantation while there was none in
the suprachoroial implantation [12]. Second, in the suprachoroidal implantation
method, the retina could be protected from heat damage generated from
electrodes. As the choroidal blood flow is much larger than in other organs in
human body, the large amount of choroidal blood flow will carry away the heat
from the electrodes and protect the retina from heat damage [23]. Third, the
surgical procedure of suprachoroidal implantation is more simple and safer than
subretinal implantation. To introduce electrodes into the suprachoroidal space,
it is not necessary to induce retinal tears and retinal detachment. In the case
of retinal detachment during subretinal implantation, the electrode could
inhibit retinal reattachement after implantation and could cause redetachement.

The suprachoroidal implantation also has weaknesses as compared to subretinal
implantation. The threshold of suprachoroidal implantation was shown to be
larger than the subretinal electrode [12]. The distance from the electrodes to
inner retinal cells might influence the impedence of electric current and this
may cause problems in making fine resolution pattern electrodes.

Our
suprachoroidal approach was similar to the previous method [20] in many
aspects. However, we placed the reference electrode on the outer scleral
surface without penetrating vitreous cavity. In our result, the placement of
the reference electrode in the extraocular space showed effective stimulation
of the retina which was evidenced by EECP and was safer than the previous
method [20] without risk of ocular damage from penetrating vitreous cavity.

In the present study, we confirmed the feasibility of our electrodes by observing
the propagation of electrically induced signals to the visual cortex by a
conventional recording method.

Moreover, the biocompatibility of the electrode was evaluated with OCT and fundus
photography. The gold electrode was safe and biocompatible as late as 16 months
after introduction. OCT gives in vivo cross-sectional retinal
images of 10 μm resolution and is an in
vivo biomicroscopy method [24, 25]. OCT was recently used to
evaluate the subretinally implanted microphotodiode arrays in cats and pigs [26, 27] and epiretinal electrodes in dogs [8]. To our knowledge, we are the first
to evaluate the biocompatibility of suprachoroidal electrodes which were
implanted into rabbits. OCT was especially
useful in detecting the subtle amount of retinal inflammation when no retinal
pathology could be found with fundus photography.

In conclusion, we have demonstrated a new design for a retinal electrical stimulator which can be used
safely in the long term. More animal experiments are needed to improve our
system and to further develop an advanced system for use in humans.

Acknowledgments

This work was supported by the Korean Science and Engineering Foundation (KOSEF)
through the Nano Bioelectronics and Systems Research Center (NBS-ERC) of Seoul
National University under Grant No. R11-2000-075-01001-0 and by the Ministry of
Health & Welfare (South Korea) through the Nano Artificial Vision Research
Center under Grant of the Korea Health 21 R&D Project (A050251). This
material was presented, in part, at The World Congress on Engineering (ICSBB
2007), July 2007.